1. Field of Invention
This invention generally relates to single photon emission computed tomography (SPECT) systems and more specifically to multi-slice systems for acquiring image data from parallel, transaxial planes in tomographic slices.
2. References
Reference is made to the following United States Letters Patent and publications:
Ter-Pogossian, Imaging Device for Computerized Tomography, U.S. Pat. No. 4,150,292 (1979).
Williams et al, Introducing SPRINT: A Single Photon Ring System for Emission Tomography, NS-29 I.E.E.E. TRANS.NUCL.SCI. 628, (1979).
Hattori et al, Emission Computed Tomograph, U.S. Pat. No. 4,389,569 (1979).
Jaszczak et al, "SPECT: Single Photon Emission Computed Tomography", NS-27 I.E.E.E. Trans.Nucl.Sci. 1137 (1980)
Hirose et al, A Hybrid Emission CT Headtome II, NS-29 I.E.E.E. Trans.Nucl.Sci. 520 (1982).
Milster et al, A Modular Scintillation Camera For Use in Nuclear Medicine, NS-31 I.E.E.E. Trans.Nucl.Sci. 578, (1984).
Wong et al, Characteristics of Small Barium Fluoride (BaF) Scintillator for High Intrinsic Resolution Time-of-Flight Positron Emission Tomography, NS-31 I.E.E.E. Trans.Nucl.Sci. 381 (1984).
Burnham et al, Design of a Cylindrical Scintillation Camera for Positron Tomographs, NS-32 I.E.E.E. Trans.Nucl.Sci. 889 (1985).
Lim et al, Triangular SPECT System for 3-D Total Organ Volume Imaging: Design Concept and Preliminary Imaging Result, NS-32 I.E.E.E. Trans.Nucl.Sci. 741 (1985)
Casey et al, A Multicrystal Two Dimensional BGO Detector System for Positron Emission Tomography, NS-33 I.E.E.E. Trans.Nucl.Sci. 460 (1986).
Chang et al, Design and Investigation of a Modular Focused Collimator Based Multiple Detector Ring System for SPECT Imaging of the Brain, 671 SPIE 200 (1986).
Genna et al, Radionuclide Annular Single Crystal Scintillator Camera with Rotating Collimator, U.S. Pat. No. 4,584,478 (1986).
Casey et al, Two Dimensional Photon Counting Position Encoder System and Process, U.S. Pat. No. 4,743,764 (1988).
Chang et al, Single Photon Emission Computed tomograph Using Focused Modular Collimators, U.S. Pat. No. 4,748,328 (1988).
Genna et al, The Development of Aspect, an Annular Single Crystal Brain Camera for High Efficiency SPECT, NS-35 I.E.E.E. Trans.Nucl.Sci. 654 (1988).
Rogers et al, SPRINT II: A Second Generation Ring-Geometry SPECT Instrument for Brain Imaging, 29 J.Nucl.Med. 760 (1988).
3. Description of Related Art
Single Photon Emission Computed Tomography (SPECT) and Positron Emission Tomography (PET) are two imaging techniques for the noninvasive imaging of a distribution of tracers in accordance with a physiological function, particularly in humans. Both systems generate tomographic images that represent the distribution of a radioisotope in one or more transaxial, and normally transverse, thin planar sections through a portion of the anatomy, such as the brain or heart. This distribution corresponds qualitatively, and in some cases quantitatively, to physiological functions being imaged.
SPECT and PET systems differ because, as their names imply, PET imaging relies upon the characteristic emission of a positron during each disintegration of a positron-emitting isotope, such as carbon-11, nitrogen-13, oxygen-15, fluorine-18, rubidium-82 or gallium-68, and the subsequent emission of two photons in opposite directions along an essentially straight line. SPECT systems rely upon the emission of a single photon that characterizes the decay of certain other isotopes, such as technetium-99m, iodine-123, thallium-201 and xenon-133.
If two photon detectors in a PET system, normally comprising scintillating crystals and photomultiplier tubes in an appropriate spatial relationship, sense the arrival of two photons within a predetermined time interval (i.e., in "coincidence"), it is assumed that a disintegration occurred in a volume between those two photon detectors. This requirement for coincidence establishes an "electronic collimation" for PET systems and eliminates the need for mechanical collimation within a transaxial plane defined by the photon detectors. The spatial resolution of PET systems in the transverse plane (i.e., a measure of the ability or uncertainty of the system to resolve small objects) is primarily dependent on the intrinsic resolution of the photon detectors. Intrinsic resolution represents a limit on the system resolving power and the possible resolution in a reconstructed image with sufficient image data. In such systems the intrinsic spatial resolution normally is stated to be a percentage (e.g., 50 to 60 percent) of the crystal size. These systems are described in the Ter-Pogossian, Wong et al, Burnham et al and Casey et al patents and publications listed above.
SPECT systems also use scintillating crystals and photomultiplier tubes as photon detectors. However, as SPECT systems rely on the emission of single photons, they can not benefit from PET's inherent electronic collimation. Rather a mechanical collimator between the photon detectors and an object being imaged limits the photons that are received by the photon detectors. More specifically, the collimator defines a number of sampling volumes through openings in the collimator that define a photon aperture. Each photon detector "sees" through such a photon aperture in the collimator only that sampling volume that lies along a corresponding sampling axis. If a photon emitted from the sampling volume travels along the sampling axis toward the collimator opening, it passes through the collimator into the corresponding photon detector. The collimator blocks photons emitted from other sites and along other directions from reaching that photon detector.
Originally discrete detector banks and conventional Anger cameras were adapted to SPECT imaging by mounting photon detectors as a head on a rotatable gantry. This gave rise to "moving detector head" systems. In such moving detector head systems, one or more detector heads including a collimator and a photon detector comprising a large flat scintillating crystal, or multiple small crystals, and photomultiplier tubes move and rotate around an object to be imaged. The photon detector localizes a scintillation event in the crystal with sufficient precision to associate each scintillation event with a specific one of the sampling volumes defined by the collimator. The head or heads rotate to sample the emitted photons from the object from many different directions. Electronics use the signals from the photomultiplier tubes to localize each scintillation event on the face of the detector head. Acquired data then is reconstructed into a tomographic image or multiple planar tomographic images each corresponding to a thin slice of the object being imaged.
More recently "fixed detector ring" systems have been developed with a stationary ring of, or cylindrical arrangement of, photon detectors in a two-dimensional form. In such systems, a rotating collimator defines, either by hardware or software means, a two-dimensional matrix of imaging positions about a cylinder while photon detectors, comprising one or more scintillating crystals and photomultiplier tubes, are disposed about the collimator's periphery. These systems define either a single transaxial planar slice or multiple, axially displaced, transaxial planar slices that tomographic image can be formed to represent the distribution of the isotope in a corresponding portion of the body.
Williams et al introduced a single-slice SPECT system of the "fixed detector ring" type. The Hattori patent also discloses a similar system with a different collimator design. In both systems, a single photon detector, including a crystal and photomultiplier tube, is associated with each imaging position in the matrix. Rogers et al and Genna et al disclose a geometry with a fixed cylindrical two-dimensional position-sensitive detector system for multi-slice imaging capability.
According to the disclosures in the Genna patent and publication, for example, a multi-slice SPECT system, also of the "fixed detector ring" type, comprises a collimator with a basically parallel hole configuration in the transaxial plane and a photon detector including a single cylindrical crystal around the collimator and large photomultiplier tubes about the periphery of the single crystal. Digital electronics apply a position determination algorithm, similar to that used in Anger logic to the information gathered by the photomultiplier tubes. These electronics localize each event to a imaging position and its associated sampling volume in a particular transaxial plane. Rogers et al describe a system comprising multiple flat modular scintillation cameras to form a polygonal cylinder for the position sensitive photon detection and a rotating slit collimator.
In all these systems each photomultiplier tube is associated with a number of imaging positions in the matrix. In such variants of gamma cameras the relationship between the spatial resolutions of the system, the collimator and the photon detectors can be given by: EQU R.sub.s.sup.2 =(R.sub.i.sup.2 +R.sub.c.sup.2) Eq. 1
where R.sub.s is the spatial resolution for the detector system, R.sub.i is the intrinsic resolution of a photon detector and R.sub.c is the collimator resolution, all in terms of the full-width at half-maximum (FWHM) of a line spread function (LSF) or the uncertainty of localization. In these gamma cameras using position sensing logic, the system resolution relies equally on both the collimator and intrinsic resolutions as defined in Eq. 1. In accordance with Eq. 1, in PET systems with no mechanical collimator, the electronic collimation basically has little uncertainty in defining the direction of photons, (i.e. as an approximation, R.sub.c .fwdarw.0), so the detector system spatial resolution is entirely dependent upon the intrinsic spatial resolution of the detectors (i.e, R.sub.s =R.sub.i).
Thus, both intrinsic and collimator resolution influence the detector system spatial resolution in each of the foregoing references. The intrinsic resolution of the photon detectors was generally comparable to or better than the resolution of the collimator over an effective collimation and SPECT imaging range (i.e., R.sub.i .ltoreq.R.sub.c). However, improvements in collimator resolution have been limited because such improvements generally are offset by decreases in system sensitivity. Thus, most of the effort for improving detector system spatial resolution has been directed to improving the intrinsic resolution of the photon detectors.
The accuracy of position determination in a position sensitive scintillating crystal system depends upon the spatial resolving capability of the detector with respect to multiple scintillation events that occur within the crystal. In gamma camera systems, each scintillation event produces responses in several photomultiplier tubes simultaneously. Over time the position determinations of those events within the crystal that occur at exactly the same interaction site have an uncertainty caused by the statistical variation of photon distribution among the photomultiplier tubes involved; that is, a map of the positions produced by the electronics will be dispersed for scintillation events actually occurring at the same site within the crystal. The degree of dispersion determines the intrinsic resolution (R.sub.i) of the photon detector system and acts to broaden the system resolution (R.sub.s) as predicted in Eq. 1.
Moreover, the uncertainty or dispersion of determined positions from adjacent imaging or collimator positions can cause overlap of data from different sites and produce significant crossover or mixing of the identifications of sampling volumes. Work to improve the intrinsic resolution has continued and continues even though it has been realized improved intrinsic resolution of a photon detector involves compromises. Specifically, improving intrinsic resolution complicates calibration procedures and correction schemes and, in some applications, can introduce limits on counting rates (i.e., the rate at which scintillation events can be recorded). If an error exists in a correction scheme or calibration procedure or a counting rate is exceeded, the resulting images could be distorted.
The Chang et al patent and publication disclose a fixed detector ring system with a single ring of collimator modules for producing a single tomographic slice. The collimator comprises a series of focused collimator modules each of which define a sampling volume; and the in-plane spatial cross-section of these sampling volumes directly determines the in-plane detector system resolution. Each collimator module also defines one imaging position in the ring. Photon detectors, comprising a single crystal and photomultiplier tube for each collimator module position, localize the collimator module and corresponding sampling volume associated with each scintillation event.
Several problems are introduced if this approach is expanded into a multiple ring configuration for imaging several planar slices simultaneously. First, the number of collimator module positions, that correspond to imaging positions in a matrix, increase proportionally with the number of rings. A corresponding increase is required in the number of photomultiplier tubes, thereby making the system more costly. Moreover, the physical size of available photomultiplier tubes limits the packing of adjacent photomultiplier tubes and has a negative impact on the overall system resolution or sensitivity. Thus, it is desirable to produce such a SPECT system in which the photomultiplier tube size is not a factor in system performance and in which photomultiplier tube costs are reduced. Moreover, it is desirable to produce a system that is adapted for implementation with other photon detector schemes.
A variety of position decoding systems have been suggested for SPECT and PET systems. In these systems the intrinsic spatial resolution of the detectors in the image plane either (1) determines detector system spatial resolution because R.sub.s =R.sub.i when R.sub.c =0 as in PET systems, or (2) strongly influences the detector spatial resolution as in prior gamma camera based SPECT systems. None of these systems suggests the use of their disclosed detector systems in a single-slice system described in the Chang et al references where the spatial resolution of the detector system depends primarily upon the spatial resolution of the collimator (i.e., R.sub.s =R.sub.c). In a design as disclosed in the Chang et al references Eq. 1 does not apply, because each detector in the transverse plane is not position sensitive. In addition, the collimator module is a discrete unit quite different from a large conventional gamma camera collimator.